Determining material stiffness using multiple aperture ultrasound

ABSTRACT

Changes in tissue stiffness have long been associated with disease. Systems and methods for determining the stiffness of tissues using ultrasonography may include a device for inducing a propagating shear wave in tissue and tracking the speed of propagation, which is directly related to tissue stiffness and density. The speed of a propagating shear wave may be detected by imaging a tissue at a high frame rate and detecting the propagating wave as a perturbance in successive image frames relative to a baseline image of the tissue in an undisturbed state. In some embodiments, sufficiently high frame rates may be achieved by using a ping-based ultrasound imaging technique in which unfocused omni-directional pings are transmitted (in an imaging plane or in a hemisphere) into a region of interest. Receiving echoes of the omnidirectional pings with multiple receive apertures allows for substantially improved lateral resolution.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a division of U.S. application Ser. No. 13/773,340,filed Feb. 21, 2013, now U.S. Pat. No. 9,339,256, which applicationclaims the benefit of U.S. Provisional Application No. 61/601,482, filedFeb. 21, 2012, all of which are incorporated by reference herein.

This application is also related to the following US patentapplications: Ser. No. 11/865,501, filed Oct. 1, 2007, now U.S. Pat. No.8,007,439, and titled “Method And Apparatus To Produce Ultrasonic ImagesUsing Multiple Apertures”; Ser. No. 12/760,375, filed Apr. 14, 2010,published as 2010/0262013 and titled “Universal Multiple ApertureMedical Ultrasound Probe”; Ser. No. 12/760,327, filed Apr. 14, 2010, nowU.S. Pat. No. 8,473,239, and titled “Multiple Aperture Ultrasound ArrayAlignment Fixture”; Ser. No. 13/279,110, filed Oct. 21, 2011, now U.S.Pat. No. 9,282,945, and titled “Calibration of Ultrasound Probes”; Ser.No. 13/272,098, filed Oct. 12, 2011 and titled “Multiple Aperture ProbeInternal Apparatus and Cable Assemblies”; Ser. No. 13/272,105, filedOct. 12, 2011, now U.S. Pat. No. 9,247,926, and titled “ConcaveUltrasound Transducers and 3D Arrays”; Ser. No. 13/029,907, filed Feb.17, 2011, now U.S. Pat. No. 9,146,313, and titled “Point SourceTransmission And Speed-Of-Sound Correction Using Multi-ApertureUltrasound Imaging”; and Ser. No. 13/690,989, filed Nov. 30, 2012 andtitled “Motion Detection Using Ping-Based and Multiple Aperture DopplerUltrasound.”

INCORPORATION BY REFERENCE

All publications and patent applications mentioned in this specificationare herein incorporated by reference to the same extent as if eachindividual publication or patent application was specifically andindividually indicated to be incorporated by reference.

FIELD

This disclosure generally relates to imaging methods and devices fordetermining a material stiffness using a multiple aperture ultrasoundprobe to produce and track ultrasonic shear waves.

BACKGROUND

Changes in tissue stiffness have long been associated with disease.Traditionally, palpation is one of the primary methods of detecting andcharacterizing tissue pathologies. It is well known that a hard masswithin an organ is often a sign of an abnormality. Several diagnosticimaging techniques have recently been developed to provide fornon-invasive characterization of tissue stiffness.

One measure of tissue stiffness is a physical quantity called Young'smodulus, which is typically expressed in units of Pascals, or morecommonly kilo Pascals (kPa). If an external uniform compression (orstress, S) is applied to a solid tissue and this induces a deformation(or strain, e) of the tissue, Young's modulus is defined simply as theratio between applied stress and the induced strain:

E=S/e.

Hard tissues have a higher Young's modulus than soft tissues. Being ableto measure the Young's modulus of a tissue helps a physician indifferentiating between benign and malignant tumors, detecting liverfibrosis and cirrhosis, detecting prostate cancer lesions, etc.

A collection of diagnostic and imaging modalities and processingtechniques have been developed to allow clinicians to evaluate tissuestiffness using ultrasonography. These techniques are collectivelyreferred to herein as Elastography. In addition to providing informationabout tissue stiffness, some elastography techniques may also be used toreveal other stiffness properties of tissue, such as axial strain,lateral strain, Poisson's Ratio, and other common strain andstrain-related parameters. Any of these or other strain-relatedparameters may be displayed in shaded grayscale or color displays toprovide visual representations of such strain-related parameters. Suchinformation may be displayed in relation to two or three dimensionaldata.

Elastography techniques may be broadly divided into two categories,“quasi-static elastography” techniques and “dynamic elastography”techniques.

In quasi-static elastography, tissue strain is induced by mechanicalcompression of a tissue region of interest, such as by pressing againsta tissue with a probe a hand or other device. In other cases, strain maybe induced by compression caused by muscular action or the movement ofadjacent organs. Images of the tissue region of interest are thenobtained in two (or more) quasi-static states, for example, nocompression and a given positive compression. Strain may be deduced fromthese two images by computing gradients of the relative local shifts ordisplacements in the images along the compression axis. Quasi-staticelastography is analogous to a physician's palpation of tissue in whichthe physician determines stiffness by pressing the tissue and detectingthe amount the tissue yields under this pressure.

In dynamic elastography, a low-frequency vibration is applied to thetissue and the speed of resulting tissue vibrations is detected. Becausethe speed of the resulting low-frequency wave is related to thestiffness of the tissue in which it travels, the stiffness of a tissuemay be approximated from wave propagation speed.

Many existing dynamic elastography techniques use ultrasound Dopplerimaging methods to detect the speed of the propagating vibrations.However, inherent limitations in standard Doppler imaging presentsubstantial challenges when attempting to measure the desiredpropagation speed. This is at least partly because the waves of mostinterest tend to have a significant propagation component in a directionperpendicular to the direction of the initial low-frequency vibration.

As used herein, the term dynamic elastography may include a wide rangeof techniques, including Acoustic Radiation Force Impulse imaging(ARFI); Virtual Touch Tissue Imaging; Shearwave Dispersion UltrasoundVibrometry (SDUV); Harmonic Motion Imaging (HMI); Supersonic ShearImaging (SSI); Spatially Modulated Ultrasound Radiation Force (SMURF)imaging.

SUMMARY OF THE DISCLOSURE

Performing Elastography with a multiple aperture ultrasound imaging(MAUI) probe provides unique advantages over prior systems and methods.For example, in some embodiments, high resolution and high frame-rateimaging capabilities of a multiple aperture probe may be combined inorder to detect a propagating shear wave as perturbations in imageframes. In other embodiments, multiple aperture Doppler imagingtechniques may be used to determine a speed of a propagating shear wave.In some embodiments, either or both of these techniques may furtherbenefit from pixel-based imaging techniques and point-sourcetransmission techniques.

In some embodiments, an ultrasound imaging system is provided,comprising a first ultrasound transducer array configured to transmit awavefront that induces a propagating shear wave in a region of interest,a second ultrasound transducer array configured to transmit circularwaveforms into the region of interest and receive echoes of the circularwaveforms, and a signal processor configured to form a plurality ofB-mode images of the region of interest from the circular waveforms at aframe rate sufficient to detect the propagating shear wave in the regionof interest.

In some embodiments, the first ultrasound transducer array comprises anarray of phased-array elements. In other embodiments, the firstultrasound transducer array comprises an annular array of piezoelectricrings, and the signal processor is further configured to focus thewavefront at various depths by adjusting phasing delays. In anotherembodiment, the first ultrasound transducer array comprises a switchedring transducer. In yet an additional embodiment, the first ultrasoundtransducer array comprises a single piezoelectric transducer.

In some embodiments, the frame rate can be at least 500 fps, at least1,000 fps, at least 2,000 fps, or at least 4,000 fps.

In one embodiment, the signal processor is further configured tocalculate a speed of the propagating shear wave by identifying a firstposition of the shear wave in a first frame of the plurality of B-modeimages, identifying a second position of the shear wave in a secondframe of the plurality of B-mode images, determining a distance traveledby the shear wave between the first frame and the second frame,determining a time elapsed between the first frame and the second frame,and dividing the distance traveled by the time elapsed.

In some embodiments, the first frame is the result of combiningsub-images formed by echoes received by multiple elements of the secondultrasound transducer array.

In another embodiment, the signal processor is configured to identifythe propagating shear wave as a point cloud moving through the region ofinterest.

In one embodiment, the signal processor is configured to define an imagewindow identifying a section of the region of interest with acombination of zooming, panning, and depth selection.

In some embodiments, the system is configured to display acontemporaneous B-mode image of a selected image window.

A method of determining a stiffness of a tissue with ultrasound isprovided, the method comprising the steps of forming a baseline image ofa region of interest with an ultrasound imaging system, transmitting anultrasonic pulse configured to induce a propagating shear wave in theregion of interest, imaging the region of interest at a frame ratesufficient to detect the propagating shear wave to form a plurality ofimage frames of the region of interest, subtracting the baseline imagefrom at least two of the formed image frames to obtain at least twodifference frames, determining a position of the propagating shear wavein the at least two difference frames, and calculating a propagationspeed of the propagating shear wave in the region of interest from thepositions in the at least two difference frames.

In some embodiments, the method further comprises calculating a tissuestiffness of the region of interest from the propagation speed.

In one embodiment, the transmitting step comprises transmitting anultrasonic pulse with a first ultrasound transducer array, and whereinthe imaging step comprises imaging the region of interest with a secondultrasound transducer array.

In another embodiment, the forming step comprises transmitting acircular waveform from a first transmit aperture and receiving echoes ona first receive aperture.

In yet another embodiment, the imaging step comprises transmitting acircular waveform from the first transmit aperture and receiving echoesof the circular waveform with the first receive aperture.

In some embodiments, the first transmit aperture and the first receiveaperture do not include overlapping transducer elements.

In another embodiment, the frame rate is at least 500 fps, at least1,000 fps, at least 2,000 fps, or at least 4,000 fps.

In some embodiments, the method further comprises identifying thepropagating shear wave as a point cloud moving through the region ofinterest.

In another embodiment, the method further comprises displaying acontemporaneous image of the region of interest, including a lineindicating a direction of transmission of the ultrasonic pulseconfigured to induce a propagating shear wave.

BRIEF DESCRIPTION OF THE DRAWINGS

The novel features of the invention are set forth with particularity inthe claims that follow. A better understanding of the features andadvantages of the present invention will be obtained by reference to thefollowing detailed description that sets forth illustrative embodiments,in which the principles of the invention are utilized, and theaccompanying drawings of which:

FIG. 1 is a schematic illustration of one embodiment of a multipleaperture ultrasound elastography probe and a propagating shear wave in aregion of interest within a viscoelastic medium.

FIG. 2 is a schematic illustration of an embodiment of a multipleaperture ultrasound elastography probe having one shear wave initiatingtransducer array and four imaging transducer arrays.

FIG. 3 is a schematic illustration of an embodiment of a multipleaperture ultrasound elastography probe having one shear wave initiatingtransducer array and two concave curved imaging transducer arrays.

FIG. 3A is an illustration of an embodiment of a multiple apertureultrasound elastography probe having a section of a continuous concavecurved array designated as a shear-wave pulse initiating area.

FIG. 3B is an illustration of an embodiment of a multiple apertureultrasound elastography probe comprising a continuous 21) concavetransducer array configured for 3D imaging with one group of elementsdesignated as a shear-wave pulse initiating area.

FIG. 4 is a schematic illustration of an annular array which may be usedfor the shear wave initiating transducer array or one or more of theimaging transducer arrays.

FIG. 5 is a flow chart illustrating one embodiment of a high resolutionmultiple aperture imaging process.

FIG. 6 is a flow chart illustrating one embodiment of a high frame ratemultiple aperture imaging process.

FIG. 7 is a flow chart illustrating one embodiment of an elastographydata capture process.

FIG. 8 is an example of a difference frame showing perturbation causedby a propagating shear wave.

DETAILED DESCRIPTION

The various embodiments will be described in detail with reference tothe accompanying drawings. References made to particular examples andimplementations are for illustrative purposes, and are not intended tolimit the scope of the invention or the claims.

In some embodiments, ultrasound imaging methods are provided in which amechanical wave having a shear component and a compression component isgenerated in a viscoelastic medium (such as biological tissue). Thespeed of the resulting shear wave propagation may be measured whileimaging the medium at a high frame-rate as the shear wave propagatesthrough the medium. Speed of the propagating shear may be determined byidentifying the changing position of the shear wave in a plurality offrames obtained at known time intervals. As will be described in furtherdetail below, various embodiments of ping-based and multiple apertureultrasound imaging are particularly well-suited to obtaining highresolution and high frame-rate images for performing accurate analysisof tissue stiffness using these methods. In some embodiments aqualitative and/or quantitative analysis of received echo data may beperformed to identify regions of different hardness as compared with therest of the viscoelastic medium.

Embodiments herein provide systems and methods for performing ultrasoundelastography to determine the shear modulus of a tissue. In someembodiments, a method of determining a shear modulus comprisestransmitting a mechanical shear wave into a test medium, then imagingthe test medium using a high frame rate B-mode ultrasound imagingtechnique as the shear wave propagates through the medium. By comparingeach image frame taken during propagation of the shear wave with areference image generated prior to transmitting the shear wave, apropagation velocity may be determined.

Although the various embodiments are described herein with reference toimaging and evaluating stiffness of various anatomic structures, it willbe understood that many of the methods and devices shown and describedherein may also be used in other applications, such as imaging andevaluating non-anatomic structures and objects. For example, theultrasound probes, systems and methods described herein may be adaptedfor use in non-destructive testing or evaluation of various mechanicalobjects, structural objects or materials, such as welds, pipes, beams,plates, pressure vessels, layered structures, soil, earth, concrete,etc. Therefore, references herein to medical or anatomic imagingtargets, tissues, or organs are provided merely as non-limiting examplesof the nearly infinite variety of targets that may be imaged orevaluated using the various apparatus and techniques described herein.

Introduction to Key Terms & Concepts

As used herein the terms “ultrasound transducer” and “transducer” maycarry their ordinary meanings as understood by those skilled in the artof ultrasound imaging technologies, and may refer without limitation toany single component capable of converting an electrical signal into anultrasonic signal and/or vice versa. For example, in some embodiments,an ultrasound transducer may comprise a piezoelectric device. In otherembodiments, ultrasound transducers may comprise capacitivemicromachined ultrasound transducers (CMUT).

Transducers are often configured in arrays of multiple individualtransducer elements. As used herein, the terms “transducer array” or“array” generally refers to a collection of transducer elements mountedto a common backing plate. Such arrays may have one dimension (1D), twodimensions (2D), 1.X dimensions (e.g., 1.5D, 1.75D, etc.) or threedimensions (3D) (such arrays may be used for imaging in 2D, 3D or 4Dimaging modes). Other dimensioned arrays as understood by those skilledin the art may also be used. Annular arrays, such as concentric circulararrays and elliptical arrays may also be used. An element of atransducer array may be the smallest discretely functional component ofan array. For example, in the case of an array of piezoelectrictransducer elements, each element may be a single piezoelectric crystalor a single machined section of a piezoelectric crystal.

As used herein, the terms “transmit element” and “receive element” maycarry their ordinary meanings as understood by those skilled in the artof ultrasound imaging technologies. The term “transmit element” mayrefer without limitation to an ultrasound transducer element which atleast momentarily performs a transmit function in which an electricalsignal is converted into an ultrasound signal. Similarly, the term“receive element” may refer without limitation to an ultrasoundtransducer element which at least momentarily performs a receivefunction in which an ultrasound signal impinging on the element isconverted into an electrical signal. Transmission of ultrasound into amedium may also be referred to herein as “insonifying.” An object orstructure which reflects ultrasound waves may be referred to as a“reflector” or a “scatterer.”

As used herein, the term “aperture” may refer to a conceptual “opening”through which ultrasound signals may be sent and/or received. In actualpractice, an aperture is simply a single transducer element or a groupof transducer elements that are collectively managed as a common groupby imaging control electronics. For example, in some embodiments anaperture may be a physical grouping of elements which may be physicallyseparated from elements of an adjacent aperture. However, adjacentapertures need not necessarily be physically separated.

It should be noted that the terms “receive aperture,” “insonifyingaperture,” and/or “transmit aperture” are used herein to mean anindividual element, a group of elements within an array, or even entirearrays with in a common housing, that perform the desired transmit orreceive function from a desired physical viewpoint or aperture. In someembodiments, such transmit and receive apertures may be created asphysically separate components with dedicated functionality. In otherembodiments, any number of send and/or receive apertures may bedynamically defined electronically as needed. In other embodiments, amultiple aperture ultrasound imaging system may use a combination ofdedicated-function and dynamic-function apertures.

As used herein, the term “total aperture” refers to the total cumulativesize of all imaging apertures. In other words, the term “total aperture”may refer to one or more dimensions defined by a maximum distancebetween the furthest-most transducer elements of any combination of sendand/or receive elements used for a particular imaging cycle. Thus, thetotal aperture is made up of any number of sub-apertures designated assend or receive apertures for a particular cycle. In the case of asingle-aperture imaging arrangement, the total aperture, sub-aperture,transmit aperture, and receive aperture will all have the samedimensions. In the case of a multiple aperture imaging arrangement, thedimensions of the total aperture includes the sum of the dimensions ofall send and receive apertures.

In some embodiments, two apertures may be located adjacent one anotheron a continuous array. In still other embodiments, two apertures mayoverlap one another on a continuous array, such that at least oneelement functions as part of two separate apertures. The location,function, number of elements and physical size of an aperture may bedefined dynamically in any manner needed for a particular application.Constraints on these parameters for a particular application will bediscussed below and/or will be clear to the skilled artisan.

Elements and arrays described herein may also be multi-function. Thatis, the designation of transducer elements or arrays as transmitters inone instance does not preclude their immediate redesignation asreceivers in the next instance. Moreover, embodiments of the controlsystem herein include the capabilities for making such designationselectronically based on user inputs, pre-set scan or resolutioncriteria, or other automatically determined criteria.

Inducing Shear Waves

The propagation velocity of shear waves in tissue is related to thestiffness (Young's modulus or shear modulus) and density of tissue bythe following equation:

E=3ρ·c ²

where c is the propagation velocity of shear wave, E is Young's modulus,and p is the tissue density. Because the density of tissues tends tovary minimally, and because the speed term is squared, elasticity may becalculated by assuming an approximate density value and measuring onlythe speed of shear wave propagation. In some cases, the assumed densityvalue may vary depending on known information about the tissue beingimaged, such as an approximate range of densities for known organtissues. For example, liver tissue may have a density of approximately1.05 kg/l, heart tissue may be about 1.03 kg/l, and skeletal muscletissue may be about 1.04 kg/l. Variations in tissue elasticity are knownto be associated with various disease states. Therefore, cancers orother pathological conditions may be detected in tissue by measuring thepropagation velocity of shear waves passing through the tissue.

In some embodiments, a shear wave may be created within tissue byapplying a strong ultrasound pulse to the tissue. In some embodiments,the shear wave generating ultrasound pulse (also referred to herein asan “initiating” pulse or an “init” pulse) may exhibit a high amplitudeand a long duration (e.g., on the order of 100 microseconds). Theultrasound pulse may generate an acoustic radiation force to push thetissue, thereby causing layers of tissue to slide along the direction ofthe ultrasound pulse. These sliding (shear) movements of tissue may beconsidered shear waves, which are of low frequencies (e.g., from 10 to500 Hz) and may propagate in a direction perpendicular to the directionof the ultrasound pulse.

Ultrasound shear waves typically result in only a few microns of tissuedisplacement. Since this amount is less than the resolution of mostimaging systems, detecting the displacement carries additionalchallenges. In some embodiments, tissue displacement induced by shearwaves may be detected in terms of the phase shift of the return ofB-mode imaging echoes.

The propagation speed of a shear wave is typically on the order of about1 to 10 m/s (corresponding to tissue elasticity from 1 to 300 kPa).Consequently, a propagating shear wave may cross a 6 cm wide ultrasoundimage plane in about 6 to 60 milliseconds. Thus, in order to collect atleast three images of a fast-moving shear waves in a 6 cm wide image, aframe rate of at least 500 frames per second may be required. Mostcurrent radiology ultrasound systems refresh a complete image only every17 to 33 milliseconds (corresponding to frame rates of about 58 to about30 frames per second), which is too slow to image a propagating shearwave because the shear wave will have disappeared from the field of viewbefore a single frame can be acquired. In order to capture shear wavesin sufficient detail, frame rates of a thousand or more images persecond are needed.

High Frame Rate Ultrasound Imaging

The frame rate of a scanline-based ultrasound imaging system is thepulse-repetition frequency (PRF, which is limited by the round-triptravel time of ultrasound in the imaged medium) divided by the number ofscanlines per frame. Typical scanline-based ultrasound imaging systemsuse between about 64 and about 192 scanlines per frame, resulting intypical frame rates of only about 50 frames per second.

By using ping-based ultrasound imaging techniques, some ultrasoundimaging systems and methods are capable of achieving frame rates on theorder of thousands of frames per second. Some embodiments of suchsystems and methods are able to obtain an entire 2D image from a singletransmit pulse, and can achieve a pulse rate (and therefore, a framerate) of 4000 per second or higher when imaging to a depth of 18 cm.With this refresh rate it is possible to capture a shear wave atincrements of about 2.5 mm of travel for the fastest waves, and evenshorter increments for slower shear waves. When imaging at shallowerdepths, even higher frame rates may be achieved. For example, whenimaging at a depth of 2 cm, a ping-based ultrasound imaging system mayachieve a pulse rate (and therefore, a frame rate) of about 75,000frames per second. Still higher frame rates may be achieved bytransmitting overlapping pulses or pings (e.g., as described below).

In contrast to conventional scanline-based phased array ultrasoundimaging systems, some embodiments of multiple aperture ultrasoundimaging systems may use point source transmission during the transmitpulse. An ultrasound wavefront transmitted from a point source (alsoreferred to herein as a “ping” or an unfocused ultrasound wavefront)illuminates the entire region of interest with each circular orspherical wavefront. Echoes received from a single ping received by asingle receive transducer element may be beamformed to form a completeimage of the insonified region of interest. Combining data and imagesfrom multiple receive transducers across a wide probe, and combiningdata from multiple pings, very high resolution images may be obtained.Moreover, such a system allows for imaging at a very high frame rate,since the frame rate is limited only by the ping repetitionfrequency—i.e., the inverse of the round-trip travel time of atransmitted wavefront travelling between a transmit transducer element,a maximum-depth reflector, and a furthest receive transducer element. Insome embodiments, the frame rate of a ping-based imaging system may beequal to the ping repetition frequency alone. In other embodiments, ifit is desired to form a frame from more than one ping, the frame rate ofa ping-based imaging system may be equal to the ping repetitionfrequency divided by the number of pings per frame.

As used herein the terms “point source transmission” and “ping” mayrefer to an introduction of transmitted ultrasound energy into a mediumfrom a single spatial location. This may be accomplished using a singleultrasound transducer element or combination of adjacent transducerelements transmitting together. A single transmission from saidelement(s) may approximate a uniform spherical wave front, or in thecase of imaging a 2D slice it creates a uniform circular wave frontwithin the 2D slice. In some cases, a single transmission of a circularor spherical wave front from a point source transmit aperture may bereferred to herein as a “ping” or a “point source pulse” or an“unfocused pulse.”

Point source transmission differs in its spatial characteristics from ascanline-based “phased array transmission” or a “directed pulsetransmission” which focuses energy in a particular direction (along ascanline) from the transducer element array. Phased array transmissionmanipulates the phase of a group of transducer elements in sequence soas to strengthen or steer an insonifying wave to a specific region ofinterest.

In some embodiments, multiple aperture imaging using a series oftransmit pings may operate by transmitting a point-source ping from afirst transmit aperture and receiving echoes of the transmitted pingwith elements of two or more receive apertures. A complete image may beformed by triangulating the position of reflectors based on delay timesbetween transmission and receiving echoes. As a result, each receiveaperture may form a complete image from echoes of each transmitted ping.In some embodiments, a single time domain frame may be formed bycombining images formed from echoes received at two or more receiveapertures from a single transmitted ping. In other embodiments, a singletime domain frame may be formed by combining images formed from echoesreceived at one or more receive apertures from two or more transmittedpings. In some such embodiments, the multiple transmitted pings mayoriginate from different transmit apertures.

“Beamforming” is generally understood to be a process by which imagingsignals received at multiple discrete receptors are combined to form acomplete coherent image. The process of ping-based beamforming isconsistent with this understanding. Embodiments of ping-basedbeamforming generally involve determining the position of reflectorscorresponding to portions of received echo data based on the path alongwhich an ultrasound signal may have traveled, an assumed-constant speedof sound and the elapsed time between a transmit ping and the time atwhich an echo is received. In other words, ping-based imaging involves acalculation of distance based on an assumed speed and a measured time.Once such a distance has been calculated, it is possible to triangulatethe possible positions of any given reflector. This distance calculationis made possible with accurate information about the relative positionsof transmit and receive transducer elements and the speed-of-ultrasoundin the imaged medium. As discussed in Applicants' previous applicationsreferenced above, multiple aperture and other probes may be calibratedto determine the acoustic position of each transducer element to atleast a desired degree of accuracy, and such element positioninformation may be digitally stored in a location accessible to theimaging or beamforming system.

FIG. 1 schematically illustrates one embodiment of a multiple apertureultrasound probe 10 configured for performing elastography. The probe 10of FIG. 1 includes two imaging transducer arrays 14, 16 and one shearwave initiating transducer array, which is referred to herein as an“init” transmit transducer array 12. An init transducer array may beconfigured for transmitting a relatively low frequency shear-waveinitiating pulse (also referred to herein as an “init pulse”).

The probe 10 may also be configured to be connected to an electroniccontroller 100 configured to electronically control transmitted andreceived ultrasonic signals. The controller may be configured totransmit phased array or ping ultrasound signals, to receive and processechoes received by the imaging transducer arrays, to perform a receivebeamforming process, and to form B-mode images from the received andprocessed echoes. The controller 100 may also be configured to controltransmission of shear wavefronts from the init array, and may beconfigured determine a position of a shear wave and an elasticity oftissue in a region of interest according to any of the embodimentsdescribed herein. The controller 100 may also be configured to controlimage formation, image processing, echo data storage, or any otherprocess, including the various methods and processes described herein.In some embodiments, some or all of the controller 100 can beincorporated into the probe. In other embodiments, the controller can beelectronically coupled to the probe (e.g., by a wired or wirelesselectronic communication method), but physically separate from the probeitself. In still further embodiments, one or more separate additionalcontrollers may be electronically connected to the probe 10 and/or tothe controller 100. Such additional controllers may be configured toexecute any one or more of the methods or processes described herein.

In the embodiment illustrated in FIG. 1, the init transducer array 12 islocated centrally in between left 14 and right 16 lateral imagingtransducer arrays. In alternative embodiments, an init array may belocated in any other position, such as the left position 14, the rightposition 16 or another position in addition to those shown in FIG. 1. Infurther embodiments, any one of several transducer arrays in a multipleaperture probe may be temporarily or permanently assigned and controlledto operate as an init array.

In further embodiments, an init transducer need not necessarily be aseparate array. Rather, a single transducer element or a group oftransducer elements that are part of a larger array that may otherwisebe used for imaging may be temporarily or permanently designated andcontrolled/operated as an init array.

As will be discussed in further detail below, the imaging transducerarrays 14, 16 of the probe 10 may be used for imaging the region ofinterest 50. The imaging transducer arrays 14, 16 may comprise anytransducer array construction suitable for ultrasound imaging, such as1D, 1.XD, 2D arrays of piezoelectric crystals or CMUT elements.

Embodiments of multiple aperture ultrasound imaging probes may includeany number of imaging apertures in a wide range of physicalarrangements. For example, FIG. 2 illustrates an embodiment of amultiple aperture elastography probe 11 comprising a central inittransducer array 12 and two pairs of imaging arrays 14, 15, 16, 17 allfour of which may be used in a multiple aperture imaging process. Insome embodiments, the init array 12 may alternatively be in the positionof any of the other arrays 14, 15, 16, 17.

In some embodiments, multiple aperture probes may have a generallyconcave tissue-engaging surface, and may include a plurality of imagingapertures. In some embodiments, each individual aperture of a multipleaperture probe may comprise a separate and distinct transducer array. Inother embodiments, individual apertures may be dynamically and/orelectronically assigned on a large continuous transducer array.

FIG. 3 illustrates an embodiment of a multiple aperture elastographyprobe comprising a central init transducer array 12 and a pair ofconcave curved lateral imaging arrays 18, 20. In some embodiments,multiple imaging apertures may be dynamically assigned on one or both ofthe concave lateral arrays 18, 20 as described in Applicants' prior U.S.patent application Ser. No. 13/272,105, now U.S. Pat. No. 9,247,926,which is incorporated herein by reference. Alternatively, each of theconcave curved lateral arrays may be treated as a separate aperture.

FIG. 3A illustrates an embodiment of a multiple aperture elastographyprobe comprising a single continuous concave curved transducer array 19.As with other embodiments discussed above, any portion of the continuouscurved array 19 may be temporarily or permanently configured,designated, and controlled/operated as an init array.

FIG. 3B illustrates an embodiment of a multiple aperture elastographyprobe comprising a 3D array 25 as described in Applicants' priorapplication Ser. No. 13/272,105, now U.S. Pat. No. 9,247,926. A group oftransducer elements 12 is shown designated as a shear wave initiatingregion. As with the above embodiments, any other region of the 3D array25 may be designated as an init region.

In some embodiments, a probe with at least three arrays may be adaptedfor elastography by replacing at least one transducer array with a lowfrequency init transducer array. In some embodiments, an init transducerarray of a multiple aperture probe may be positioned between at leasttwo other arrays. Such probe configurations may include adjustableprobes, cardiac probes, universal probes, intravenous ultrasound (IVUS)probes, endo-vaginal probes, endo-rectal probes, transesophageal probesor other probes configured for a particular application.

Similarly, any other multiple aperture or single-aperture ultrasoundimaging probe may be adapted for use with the elastography systems andmethods described herein. In still further embodiments, an init arraymay be provided on a separate probe entirely independent of an imagingprobe. For example, an init probe may be provided with a separatehousing from the housing of the imaging probe. In some embodiments, anindependent init probe may be configured to be temporarily attached toan imaging probe. In such embodiments, such a separate init probe may becontrolled by the same ultrasound imaging system as an imaging probe, orthe init probe may be controlled independently of the imaging system. Anindependently-controlled elastography init pulse controller may besynchronized with an ultrasound imaging system in order to provide theimaging system with accurate timing information indicating the time atwhich an init pulse is transmitted.

In alternative embodiments, similar frame rates may be achieved bytransmitting a plane wave front (e.g., by transmitting simultaneouspulses from several transducers in a common array), receiving echoes,and mapping the received echoes to pixel locations using techniquessimilar to those described above. Some embodiments of such plane-wavetransmitting systems may achieve frame rates similar to those achievedwith ping-based imaging techniques.

Embodiments of Shear-Wave Initiating Transducers

Regardless of probe construction, embodiments of an init array 12 may beconfigured to transmit shear-wave initiating ultrasound pulses withfrequencies between about 1 MHz and about 10 MHz. In other embodiments,the init array 12 may be configured to transmit shear-wave initiatingultrasound pulses with a frequency up to about 18 MHz or higher. In someembodiments, an ultrasound frequency for producing init pulses may beabout half of an ultrasound frequency used for imaging. Depending onmaterials and construction, a single transducer array may be capable ofproducing both low frequency ultrasound pulses for an init pulse andrelatively high frequency ultrasound pulses for imaging. However, insome embodiments it may be desirable to use transducers optimized for arelatively narrow frequency range to allow for more efficient control ofan init pulse or an imaging pulse.

Thus, in some embodiments, an init transducer array 12 may comprise aseparate array configured to function exclusively as an init array, suchas by being optimized to function efficiently within an expected initfrequency range. As a result, in some embodiments an init array may bestructurally different than separate imaging arrays. In otherembodiments an init array may be physically identical to an imagingarray, and may differ only in terms of its operation and use.

In some embodiments, the init transducer array 12 may comprise arectangular or otherwise shaped array (e.g., a 1D, 1.xD, 2D or otherrectangular array) of piezoelectric elements. In other embodiments, theinit transducer array 12 may comprise a rectangular or otherwise shapedarray of capacitive micro-machined ultrasound transducer (CMUT)elements.

In other embodiments, the init array 12 may comprise an annular array 30as shown for example in FIG. 4. An annular array may comprise aplurality of transducer elements arranged in concentric circular orelliptical patterns. Such annular arrays 20 may also use any suitabletransducer material. In some embodiments, an init array 12 may comprisea switched ring annular transducer array.

In some embodiments, a switched-ring annular array may include adish-shaped ultrasonic transducer (e.g., a segment of a sphere) whichmay be divided into a plurality of concentric annular transducerelements of which the innermost element may be either a planar annulusor a complete dish. In some embodiments, the curvature of the frontsurface of the annular array 20 and any lens or impedance matching layerbetween the transducer and the region of interest surface may at leastpartially determine the focal length of the transducer. In otherembodiments, an annular array may be substantially planar and anacoustic lens may be employed to focus the transmitted ultrasoundenergy.

An annular array 20 may include any number of rings, such as three ringsin addition to the center disc as shown in FIG. 4. In other embodiments,an annular array may include 2, 4, 5, 6. 7. 8, 9, 10 or more rings inaddition to a center disc or dish. In some embodiments, the rings may befurther decoupled by etching, scribing, complete cutting or otherwisedividing the rings into a plurality of ring elements within each ring.In some embodiments, an annular array transducer for operating to depthsof 25 cm may have a diameter of 40 mm with the outer ring may have awidth of approximately 1.85 mm, providing a surface area of 222 mm²; theinner ring may have a width of approximately 0.8 mm and lying at anapproximate radius of 10.6 mm to provide a surface area of 55 mm².

In some embodiments, each ring (or each ring element within a ring) mayhave individual electrical connections such that each ring (or ringelement) may be individually controlled as a separate transducer elementby the control system such that the rings may be phased so as to directa shear-wave initiating pulse to a desired depth within the region ofinterest. The amplitude of the energy applied may determine theamplitude of the emitted ultrasonic waves which travel away from theface of the annular array 20.

In some embodiments the size and/or number of elements in an init arraymay be determined by the shape or other properties of the shear waves tobe produced.

In some embodiments, a shear-wave initiating pulse produced by an inittransducer array 12 may be focused during transmission to providemaximum power at the region of interest. In some embodiments, the initpulse may be focused on an init line 22 (e.g., as shown in FIGS. 1, 2and 3). The init pulse may further be focused at a desired depth toproduce a maximum disruptive power at the desired depth. In someembodiments, the axial focus line and the focused depth point may bedetermined by transmitting pulses from a plurality of transducerelements at a set of suitable delays (i.e., using “phased array”techniques). In some embodiments, transmit delays may be omitted whenusing an annular array with a series of switched rings as discussedabove.

In some embodiments, the init pulse need not be electronicallysteerable. In such embodiments, the probe may be configured to alwaystransmit an init pulse along a consistent line relative to the probe. Insome embodiments, the expected line of the init pulse may be displayedon the ultrasound display (e.g., overlaying a contemporaneous B-modeimage of the region of interest) so as to provide an operator with avisual indication of the path of the init pulse relative to the imagedregion of interest. In such embodiments, a sonographer may manipulatethe probe until the display shows a representative init line passingthrough an object to be evaluated by elastography.

In alternative embodiments, an init pulse may be electronically steeredin a direction indicated by an operator. In such embodiments, the lineof the init pulse may be selected by an operator through any appropriateuser interface interaction without the need to move the probe. In someembodiments, the user interface interaction may include a visual displayof the init line on a display screen (e.g., overlaying a contemporaneousB-mode image of the region of interest). Once a desired init pulsedirection is chosen, an init pulse may be electronically steered so asto travel along the selected line.

Embodiments for Detecting Shear Wave Propagation Rate

Returning to FIG. 1, an example of shear wave propagation will bedescribed. A shear wave may be initiated in a region of interest 50 froman init pulse from a multiple aperture elastography probe 10 (or anyother suitably configured elastography probe). As discussed above, theinit pulse may be focused along a line 22 extending from the inittransducer array 12 into the region of interest to at least a desireddepth. In some embodiments, the line 22 may be perpendicular to the inittransducer array 12. An initial pulse 52 transmitted along the init line22 will tend to induce a wave front 56 propagating outwards from theline 22 within the image plane. The propagating wavefront 56 induced bythe init pulse will push the tissue in the direction of propagation. Anelastic medium such as human tissue will react to this push by arestoring force which induces mechanical waves including shear waveswhich propagate transversely from the line 22 in the tissue.

Embodiments of elastographic imaging processes will now be describedwith reference to the probe construction of FIG. 1 and the flow chartsof FIGS. 5-7. These processes may be used with any suitably configuredprobe as described above. In some embodiments, the left and rightlateral transducer arrays 14, 16 may be used to image the region ofinterest 50 with either, both or a combination of a high frame rateultrasound imaging technique and a high resolution multiple apertureultrasound imaging technique. These techniques are summarized below, andfurther details of these techniques are provided in U.S. patentapplication Ser. No. 13/029,907, now U.S. Pat. No. 9,146,313, whichillustrates embodiments of imaging techniques using transmission of acircular wavefront and using receive-only beamforming to produce anentire image from each pulse or “ping” (also referred to as ping-basedimaging techniques).

The terms “high resolution imaging” and “high frame rate imaging” areused herein as abbreviated names for alternative imaging processes.These terms are not intended to be limiting or exclusive, as the “highresolution imaging” process may also be operated at a high frame raterelative to other imaging techniques, and the “high frame rate imaging”process may also produce images of a substantially higher resolutionthan other imaging techniques. Furthermore, the rate of shear wavepropagation may be detected using high frame rate imaging techniquesand/or high resolution imaging techniques other than those described orreferenced herein.

FIG. 5 illustrates an embodiment of a high resolution multiple apertureimaging process 60 that may use a multiple aperture ultrasound imagingprobe such as that shown in FIG. 1. In some embodiments, one or both ofthe imaging arrays 14, 16 may include one or more transducer elementstemporarily or permanently designated as transmit elements T1 throughTn. The remaining transducer elements of one or both of the imagingarrays 14, 16 may be designated as receive elements.

In some embodiments, a high resolution multiple aperture ultrasoundimaging process 60 may comprise transmitting a series of successivepulses from a series of different transmit apertures (T1 . . . Tn) 62,receiving echoes 64 from each pulse with a plurality of elements on areceive aperture, and obtaining a complete image 66 from echoes receivedfrom each transmit pulse. These images may then be combined 68 into afinal high-resolution image. Embodiments of such a high resolutionmultiple aperture imaging process may be substantially similar to theprocess shown and described in Applicants' prior U.S. patent applicationSer. No. 13/029,907, now U.S. Pat. No. 9,146,313, referenced above.

As indicated in FIG. 5, during a first cycle of a high resolutionimaging process, the steps of transmitting an ultrasound signal 62A,receiving echoes 64A, and forming an image 66A may be performed using afirst transmit transducer T1. During a second cycle, signals may betransmitted 62B from a different transmit transducer Ti, echoes may bereceived 64B, and a second image may be formed 66B. The process of steps62 x-66 x may be repeated using n different transmit transducers whichmay respectively be located at any desired position within an ultrasoundprobe. Once a desired number of image (also referred to as image layers)have been formed, such image layers may be combined 68 into a singleimage frame, thereby improving image quality. If desired, the process 60may then be repeated to obtain multiple time-domain frames which maythen be consecutively displayed to a user.

FIG. 6 illustrates an embodiment of a high frame rate imaging process70. In some embodiments, a high frame rate ultrasound imaging process 70may comprise transmitting successive pings from a single transmitaperture Tx 72, forming a complete image 76 from echoes received 74 fromeach transmitted ping 72, and treating each image 76 as a successivetime domain frame. In this way, slight changes in the position ofreflectors in the region of interest 50 can be sampled at a very highframe rate.

As indicated in FIG. 6, during a first cycle, a ping may be transmittedfrom a chosen transmit transducer Tx 72A, echoes may be received 74A anda first frame may be formed 76A. The same cycle of steps transmitting72B and receiving 74B may then be repeated to produce a second frame76B, a third frame (steps 72C, 74C, 76C), and as many subsequent framesas desired or needed as described elsewhere herein.

In some embodiments, a maximum frame rate of an imaging system usingping-based imaging techniques may be reached when a ping repetitionfrequency (i.e., the frequency at which successive pings aretransmitted) is equal to an inverse of the round trip travel time (i.e.,the time for an ultrasound wave to travel from a transmit transducer toa reflector at a desired distance from the transducer, plus the time foran echo to return from the reflector to a receive transducer along thesame or a different path). In other embodiments, overlapping pings maybe used with coded excitation or other methods of distinguishingoverlapping echoes. That is, a second ping may be transmitted before allechoes from a first ping are received. This is possible as long as thetransmitted ping signals may be coded or otherwise distinguished suchthat echoes of a first ping may be recognized as distinct from echoes ofa second ping. Several coded excitation techniques are known to thoseskilled in the art, any of which may be used with a point-sourcemultiple aperture imaging probe. Alternatively, overlapping pings mayalso be distinguished by transmitting pings at different frequencies orusing any other suitable techniques. Using overlapping pings, evenhigher imaging frame rates may be achieved.

In some embodiments, prior to initiating an elastographic imagingprocess, an imaging window may be defined during a B-mode imagingprocess. The defined image window may be a section of the region ofinterest in which elastography is to be performed. For example, theimage window may be defined after any combination of probe positioning,depth-selection, zooming, panning, etc. In some embodiments, an imagewindow may be as large as an entire insonified region of interest. Inother embodiments, an image window may be only a smaller section of thecomplete region of interest (e.g., a “zoomed-in” section). In someembodiments, an image window may be defined after an imaging sessionusing echo data retrieved from a raw data memory device.

FIG. 7 illustrates an embodiment of an elastography process 80 using aprobe such as that shown in FIG. 1. In the illustrated embodiment, anelastography process 80 may generally involve the steps of obtaining 82and storing 84 a baseline image, transmitting a shear-wave initiatingpulse (an init pulse) 86 into the region of interest 50, imaging theregion of interest 50 using a high frame rate imaging process 88, andsubtracting the baseline image 90 from each frame obtained during thehigh frame rate imaging process 88. The remaining series of “differenceframes” can then be analyzed to obtain information about the tissuedisplaced by the shear wave 56 propagating through the tissue of theregion of interest 50. The propagation speed of the shear wave 56 may beobtained through analysis of the perturbation of tissue in thetime-series of difference frames.

In some embodiments, while imaging a selected image window within aregion of interest with an elastography-enabled ultrasound probe, aninit line 22 (shown in FIG. 1) may be displayed on an ultrasound imagedisplay screen overlying an image of the target region. In someembodiments, the ultrasound imaging system may continuously image theregion of interest with a high resolution imaging process as discussedabove with reference to FIG. 5. Alternatively, any other desiredultrasound imaging process may be used to obtain an image of the regionto be analyzed by an elastography process.

Once the probe 10 is in a desired orientation such that the init line 22intersects a desired target object or portion of the region of interest,an elastography depth may be selected, and an elastography process 80may be initiated. In some embodiments, an elastography depth may beselected by an operator via a suitable user interface action. In otherembodiments, an elastography depth may be selected automatically by anultrasound imaging control system. In some embodiments, an elastographyprocess may be initiated manually by an operator of the ultrasoundsystem. In other embodiments, an elastography process 80 may beinitiated automatically by an ultrasound system upon automaticidentification of a structure to be inspected.

As shown in the embodiment of FIG. 7, an elastography process 80 using aprobe such as that shown in FIG. 1 (or any other suitably configuredprobe) may begin by obtaining 82 and storing 84 a baseline image of thetarget region of interest 50. In one embodiment, the baseline image maybe formed by obtaining a single frame using a high-frame-rate imagingprocess such as that described above. In such embodiments, a baselineimage may be formed by transmitting an imaging pulse from a singletransducer element Tx from a first of the lateral transducer arrays 14,16 (e.g., the right array 16), and receiving echoes on multiple elementsof the second of the lateral transducer arrays 14, 16 (e.g., the leftarray 14). In some embodiments, echoes from the transmit pulse may alsobe received by receive elements on the first transducer array (e.g. theright array 16). The baseline image may then be formed and stored 84 foruse in subsequent steps. In an alternative embodiment, the baselineimage may be obtained 82 using a high resolution imaging process such asthat described above.

After obtaining a baseline image 82, the init transducer array may beoperated to transmit a shear-wave initiating pulse 86 into the region ofinterest. An init pulse may be produced by any suitable devices andmethods as described above. In some embodiments, the shear waveinitiating pulse may be focused along a displayed init line 22, and maybe focused at a particular depth within the region of interest.

After an init pulse is transmitted 86, the system may begin imaging theregion of interest at a high frame rate 88 using the lateral imagingarrays 14, 16. In some embodiments, the high frame rate imaging processmay comprise the process described above with reference to FIG. 6. Inone embodiment, the high frame rate imaging process may comprisetransmitting a series of transmit pulses from a single transmit apertureTx, and receiving echoes at a plurality of elements on at least onereceive aperture. In some embodiments, the high frame rate imaging 88may be performed by transmitting ultrasound pulses from the sametransmit element (or aperture) as that used in the step of obtaining abaseline image 82. In some embodiments, the high frame rate imaging maycontinue at least until propagation of the induced shear wave hasstopped or has progressed to a desired degree. A duration of highframe-rate imaging time may be calculated in advance based on anexpected minimum propagation speed and an image size. Alternatively, thehigh frame rate imaging 88 may be stopped upon detecting the shearwave's propagation at an extent of an imaging frame.

In some embodiments, forming a single frame during a high frame rateimaging process 88 may include combining image layers obtained fromechoes received at different receiving transducer elements. For example,separate images may be formed from echoes received by each individualtransducer element of a receive aperture to form a single improvedimage. Then, a first image produced by echoes received by all elementsof a first receive aperture may be combined with a second image producedby echoes received by all elements of a second receive aperture in orderto further improve the quality of the resulting image. In someembodiments, the image resulting from such combinations may then be usedas a single frame in the high frame rate imaging process 88. Furtherexamples of such image combining are described in U.S. patentapplication Ser. No. 13/029,907, now U.S. Pat. No. 9,146,313, referencedabove.

In some embodiments, the baseline image may then be subtracted 90 fromeach individual frame obtained in the high frame rate imaging process88. For example, each pixel value of a single frame may be subtractedfrom the value of each corresponding pixel in the baseline image. Theimage resulting from such subtraction may be referred to as a“difference image” or a “difference frame.” The difference images thusobtained will include pixel values representing substantially only theshear waveform plus any noise.

In some embodiments, the steps of obtaining a baseline image 82,transmitting an init pulse 86 continuously imaging at a high frame rate88, and obtaining difference image frames 90 may be repeated as manytimes as desired. The difference images from such multiple cycles may beaveraged or otherwise combined in order to improve a signal to noiselevel.

The propagating shear waveform may be detected along lines transverse tothe direction of the init pulse (e.g., as shown in FIG. 1) by detectingperturbation (i.e., small changes in an otherwise ‘normal’ pattern) insubsequent difference frames. The speed of the shear wave's propagationmay be obtained by determining the position of the shear wave inmultiple image frames obtained at known time intervals.

In some cases, the perturbation caused by a propagating shear wave mayproduce a relatively disbursed image of the propagating wave front. Forexample, perturbation may appear in a difference frame as a specklepattern 92 such as that shown in FIG. 8. An approximate center line 94of the point cloud 92 may be determined and treated as representative ofthe position of the propagating shear wavefront. In some embodiments, aline, curve or other path 94 may be fit to the point cloud 92 using anysuitable path fit algorithm. For example, in some embodiments anabsolute value of the difference frame may be calculated, and a localposition of the shear wave may be determined by averaging the positionof the nearest x points.

In some embodiments, the analysis may be limited to only a portion ofthe point cloud 92 (and/or a corresponding center line 94). For example,if it is determined (by visual inspection or by automated analysis) thata small segment of the shear wavefront is propagating faster thanadjacent segments, the region(s) of apparent higher or lower propagationspeed may be selected, and the speed of propagation may be calculatedfor only that portion of the shear wavefront.

By calculating a distance between the focused init line 22 and the fitline 94 in a given difference frame, an approximate position of theshear wave in the given difference frame may be calculated. The rate ofpropagation of the wavefront between any two frames may be determined bydividing the distance traveled by the shear wave by the time thatelapsed between obtaining the two frames. In alternative embodiments,the position of a shear wave in any given frame may be measured relativeto any other suitable datum.

In various embodiments, the number of frames needed to measure thepropagation speed of a shear wave may vary. In some embodiments anapproximate speed measurement may be obtained from as few as two orthree frames obtained at known time intervals. In other embodiments, atleast ten frames obtained at known time intervals may be needed toobtain a sufficiently accurate time measurement. In further embodiments,at least 100 frames obtained at known time intervals may be used toobtain a more accurate time measurement. In still further embodiments,200 frames or more may be used. Generally, the accuracy of shear wavepropagation speed measurements may increase with the number of framesfrom which such measurements are made. As the number of framesincreases, so does computational complexity, so the number of frames tobe used may be balanced with available processing capabilities.

When more than two frames are available to be used for measuringpropagation speed, any number of algorithms may be used. For example, insome embodiments the shear wave position may be detected in eachavailable frame, a speed may be calculated between each consecutive pairof frames, and the results of all such speed measurements may beaveraged to obtain a single speed value. In other embodiments, speedmeasurements may be calculated based on time intervals and relativeshear wave positions between different and/or variable numbers offrames. For example, propagation speed may be calculated between everythree frames, every five frames, every 10 frames, every 50 frames, etc.Such measurements may then be averaged with one another and/or withmeasurements obtained from consecutive frame pairs. Weighted averagesmay also be used in some embodiments.

In some embodiments, an entire elastography process 80 (FIG. 7) may berepeated at different focus depths relative to the init transducer array12. In some embodiments, un-beamformed elastography echo data obtainedat various depths may be stored and combined into a single 2D or 3D dataset for further post processing and/or for later viewing and analysis.In various embodiments, un-beamformed elastography echo data may becaptured and stored for later processing on the imaging system or anyother suitable computing hardware.

In alternative embodiments, the propagation speed of a shear wave may bemeasured by detecting the speed of moving/displaced tissues using themultiple aperture Doppler techniques described in Applicant's co-pendingU.S. patent application Ser. No. 13/690,989, filed Nov. 30, 2012, titled“Motion Detection Using Ping-Based And Multiple Aperture DopplerUltrasound.”

Once the shear wave is captured and its propagation speed is measured,the hardness of the tissue in the region of interest, as quantified byYoung's modulus (E) can be measured or determined by a controller,signal processor or computer. Elasticity (E) and shear wave propagationspeed (c) are directly related through the simple formula:

E=3ρc ²

Where p is the density of tissue expressed in kg/m³. Because the densityof tissues tends to vary minimally, an approximate density value may beassumed for the purpose of calculating elasticity using a measuredpropagation speed value. The fact that the speed term is squared furtherminimizes the effect of any error in the assumed density value. Thus,the elasticity of the tissue may be calculated after measuring only theshear wave propagation velocity c and using an assumed approximate valuefor tissue density.

In some embodiments, the density value may be stored in a digital memorydevice within or electronically accessible by the controller. In otherembodiments, the density value may be manually entered or edited by auser via any suitable user interface device. Once the speed of shearwave propagation has been measured for a desired area within the regionof interest, the controller may retrieve the density value and calculatethe elasticity for the desired area.

In some embodiments, elasticity estimates may be overlaid on an image ofthe region of interest. In some embodiments, such an overlay may beprovided as a color coded shaded image, showing areas of high elasticityin contrasting colors to areas of relatively low elasticity.Alternatively, a propagating shear wave may be displayed on an image. Insome embodiments, a propagating shear wave may be displayed as ananimated moving line, as changing colors, as a moving point cloud or inother ways. In further embodiments, a numeric value of a shear wavepropagation speed may be displayed. In other embodiments, numeric valuesof elasticity may be displayed on an image of the region of interest.Soft tissues will tend to have relatively small values of elasticity,and liquid-filled areas do not conduct shear waves at all.

Raw Echo Data Memory

Various embodiments of the systems and methods described above may befurther enhanced by using an ultrasound imaging system configured tostore digitized echo waveforms during an imaging session. Such digitalecho data may be subsequently processed on an imaging system or on anindependent computer or other workstation configured to beamform andprocess the echo data to form images. In some embodiments, such aworkstation device may comprise any digital processing system withsoftware for dynamically beamforming and processing echo data using anyof the techniques described above. For example, such processing may beperformed using data processing hardware that is entirely independent ofan ultrasound imaging system used to transmit and receive ultrasoundsignals. Such alternative processing hardware may comprise a desktopcomputer, a tablet computer, a laptop computer, a smartphone, a serveror any other general purpose data processing hardware.

In various embodiments, received echo data (including echoes receivedduring a high frame rate imaging process) may be stored at variousstages from pure analog echo signals to fully processed digital imagesor even digital video. For example, a purely raw analog signal may bestored using an analog recording medium such as analog magnetic tape. Ata slightly higher level of processing, digital data may be storedimmediately after passing the analog signal through an analog-to-digitalconverter. Further processing, such as band-pass filtering,interpolation, down-sampling, up-sampling, other filtering, etc. may beperformed on the digitized echo data, and raw data may be stored aftersuch additional filtering or processing steps. Such raw data may then bebeamformed to determine a pixel location for each received echo, therebyforming an image. Individual images may be combined as frames to formvideo. In some embodiments, it may be desirable to store digitized echodata after performing very little processing (e.g., after some filteringand conditioning of digital echo data, but before performing anybeamforming or image processing). Some ultrasound systems storebeamformed echo data or fully processed image data. Nonetheless, as usedherein, the phrases “raw echo data” and “raw data” may refer to storedecho information describing received ultrasound echoes (RX data) at anylevel of processing prior to beamforming. Raw echo data may include echodata resulting from B-mode pings, Doppler pings, or any other ultrasoundtransmit signal.

In addition to received echo data, it may also be desirable to storeinformation about one or more ultrasound transmit signals that generateda particular set of echo data. For example, when imaging with a multipleaperture ping ultrasound method as described above, it is desirable toknow information about a transmitted ping that produced a particular setof echoes. Such information may include the identity and/or position ofone or more a transmit elements as well as a frequency, magnitude, pulselength, duration or other information describing a transmittedultrasound signal. Transmit data is collectively referred herein to as“TX data”.

In some embodiments, TX data may also include information defining aline along which a shear-wave initiating pulse is transmitted, andtiming information indicating a time at which such a shear-waveinitiating pulse is transmitted relative to received echo data.

In some embodiments, such TX data may be stored explicitly in the sameraw data memory device in which raw echo data is stored. For example, TXdata describing a transmitted signal may be stored as a header before oras a footer after a set of raw echo data generated by the transmittedsignal.

In other embodiments, TX data may be stored explicitly in a separatememory device that is also accessible to a system performing abeamforming process. In embodiments in which transmit data is storedexplicitly, the phrases “raw echo data” or “raw data” may also includesuch explicitly stored TX data. In still further embodiments, transducerelement position information may be explicitly stored in the same or aseparate memory device. Such element position data may be referred to as“calibration data” or “element position data”, and in some embodimentsmay be generally included within “raw data.”

TX data may also be stored implicitly. For example, if an imaging systemis configured to transmit consistently defined ultrasound signals (e.g.,consistent magnitude, shape, frequency, duration, etc.) in a consistentor known sequence, then such information may be assumed during abeamforming process. In such cases, the only information that needs tobe associated with the echo data is the position (or identity) of thetransmit transducer(s). In some embodiments, such information may beimplicitly obtained based on the organization of raw echo data in a rawdata memory. For example, a system may be configured to store a fixednumber of echo records following each ping. In such embodiments, echoesfrom a first ping may be stored at memory positions 0 through ‘n’ (where‘n’ is the number of records stored for each ping), and echoes from asecond ping may be stored at memory positions n+1 through 2n+1. In otherembodiments, one or more empty records may be left in between echo sets.In some embodiments received echo data may be stored using variousmemory interleaving techniques to imply a relationship between atransmitted ping and a received echo data point (or a group of echoes).Similarly, assuming data is sampled at a consistent, known samplingrate, the time at which each echo data point was received may beinferred from the position of that data point in memory. In someembodiments, the same techniques may also be used to implicitly storedata from multiple receive channels in a single raw data memory device.

In some embodiments, raw TX data and raw echo data may be captured andstored during an imaging session in which an elastography process isperformed. Such data may then be later retrieved from the memory device,and beamforming, image processing, and shear-wave speed measurementsteps may be repeated using different assumptions, inputs or algorithmsin order to further improve results. For example, during suchre-processing of stored data, assumed values of tissue density orspeed-of-sound may be used. Beamforming, image layer combining, or speedmeasurement averaging algorithms may also be modified during suchre-processing relative to a real-time imaging session. In someembodiments, while reprocessing stored data, assumed constants andalgorithms may be modified iteratively in order to identify an optimumset of parameters for a particular set of echo data.

Although this invention has been disclosed in the context of certainpreferred embodiments and examples, it will be understood by thoseskilled in the art that the present invention extends beyond thespecifically disclosed embodiments to other alternative embodimentsand/or uses of the invention and obvious modifications and equivalentsthereof. Various modifications to the above embodiments will be readilyapparent to those skilled in the art, and the generic principles definedherein may be applied to other embodiments without departing from thespirit or scope of the invention. Thus, it is intended that the scope ofthe present invention herein disclosed should not be limited by theparticular disclosed embodiments described above, but should bedetermined only by a fair reading of the claims that follow.

In particular, materials and manufacturing techniques may be employed aswithin the level of those with skill in the relevant art. Furthermore,reference to a singular item, includes the possibility that there areplural of the same items present. More specifically, as used herein andin the appended claims, the singular forms “a,” “and,” “said,” and “the”include plural referents unless the context clearly dictates otherwise.As used herein, unless explicitly stated otherwise, the term “or” isinclusive of all presented alternatives, and means essentially the sameas the commonly used phrase “and/or.” Thus, for example the phrase “A orB may be blue” may mean any of the following: A alone is blue, B aloneis blue, both A and B are blue, and A, B and C are blue. It is furthernoted that the claims may be drafted to exclude any optional element. Assuch, this statement is intended to serve as antecedent basis for use ofsuch exclusive terminology as “solely,” “only” and the like inconnection with the recitation of claim elements, or use of a “negative”limitation. Unless defined otherwise herein, all technical andscientific terms used herein have the same meaning as commonlyunderstood by one of ordinary skill in the art to which this inventionbelongs.

What is claimed is:
 1. An ultrasound imaging system, comprising: a firstultrasound transducer array configured to transmit an init wavefrontthat induces a propagating shear wave in a region of interest; a secondultrasound transducer array configured to transmit circular waveformsinto the region of interest and receive echoes of the circularwaveforms; and a processor configured to form a plurality of B-modeimages of the region of interest from the circular waveforms at a framerate sufficient to detect the propagating shear wave in the region ofinterest.
 2. The system of claim 1, wherein the first ultrasoundtransducer array comprises an array of phased-array elements.
 3. Thesystem of claim 1, wherein the first ultrasound transducer arraycomprises an annular array of piezoelectric rings, and the processor isfurther configured to focus the init wavefront at various depths byadjusting phasing delays.
 4. The system of claim 3, wherein the firstultrasound transducer array comprises a switched ring transducer.
 5. Thesystem of claim 1 wherein the first ultrasound transducer arraycomprises a single piezoelectric transducer.
 6. The system of claim 1,wherein the frame rate is at least 500 fps.
 7. The system of claim 1,wherein the frame rate is at least 1,000 fps.
 8. The system of claim 1,wherein the frame rate is at least 2,000 fps.
 9. The system of claim 1,wherein the frame rate is at least 4,000 fps.
 10. The system of claim 1,wherein the processor is further configured to calculate a speed of thepropagating shear wave by identifying a first position of the shear wavein a first frame of the plurality of B-mode images, identifying a secondposition of the shear wave in a second frame of the plurality of B-modeimages, determining a distance traveled by the shear wave between thefirst frame and the second frame, determining a time elapsed between thefirst frame and the second frame, and dividing the distance traveled bythe time elapsed.
 11. The system of claim 10, wherein the first frame isthe result of combining sub-images formed by echoes received by multipleelements of the second ultrasound transducer array.
 12. The system ofclaim 11, wherein the processor is configured to form the first frameby: transmitting a first unfocused ultrasound ping from a firsttransmitter transducer element; forming a first baseline image layerwith an electronic controller using echoes of the first ultrasound pingreceived by elements of a first receive aperture; forming a secondbaseline image layer with the electronic controller using echoes of thefirst ultrasound ping received by elements of a second receive aperture;transmitting a second unfocused ultrasound ping from a secondtransmitter transducer element different than the first transmitterelement; forming a third baseline image layer with the electroniccontroller using echoes of the second ultrasound ping received byelements of the first receive aperture; forming a fourth baseline imagelayer with the electronic controller using echoes of the secondultrasound ping received by elements of the second receive aperture; andcombining the first baseline image layer, the second baseline imagelayer, the third baseline image layer, and the fourth baseline imagelayer with the electronic controller to form the baseline image.
 13. Thesystem of claim 11, wherein the processor is configured to form thesecond frame by: transmitting a first unfocused ultrasound ping from afirst transmitter transducer element of the second ultrasound transducerarray; forming a first image layer using echoes of only the firstultrasound ping received by elements of a first receive aperture of thesecond ultrasound transducer array; forming a second image layer usingechoes of only the first ultrasound ping received by elements of asecond receive aperture of the second ultrasound transducer array; andcombining the first image layer and the second image layer with theelectronic controller to form the first frame.
 14. The system of claim1, wherein the processor is configured to identify the propagating shearwave as a speckle pattern moving through the region of interest.
 15. Thesystem of claim 1, wherein the processor is configured to define animage window identifying a section of the region of interest with acombination of zooming, panning, and depth selection.
 16. The system ofclaim 15, wherein the system is configured to display a contemporaneousB-mode image of a selected image window while calculating and displayinga modulus of elasticity of an object in the region of interest.